RU2603051C2 - Adjusting measurements of the effects of acoustic radiation force for background motion effects - Google Patents

Abstract

FIELD: medicine.

SUBSTANCE: group of inventions refers to medical equipment, namely to ultrasonic diagnostic imaging systems. Ultrasonic diagnostic imaging system for shear wave analysis comprising an ultrasonic array probe containing two dimensional array of transducer elements, which transmits a push pulse along a predetermined vector to generate a shear wave, transmits tracking pulses along tracking lines adjacent to the push pulse vector, and receives echo signals from points along the tracking lines, wherein the focused push pulse has measures of height, and of the azimuth, memory for storing the tracking line echo data, motion detector responsive to the tracking line data for detecting a shear wave passing through the tracking line locations, and a display for displaying a characteristic of a detected shear wave, wherein the ultrasound array probe is further operable to transmit along background motion tracking lines, along which background motion echo signals are received, in the vicinity of the push pulse vector at different times, which are compared to sense background motion in the vicinity of a shear wave. Method for adjusting the measurement of a shear wave characteristic measured in a region of tissue, for the effect of relative motion between the tissue and an ultrasound probe, comprises the steps wherein the ultrasonic probe is used to detect a shear wave in the region of tissue, transmitting background motion echo signals along background motion tracking lines beyond the depth of the push pulse, comparing the echo signals acquired at different times to adjust measured shear wave characteristics, that is corrected due to the relative motion.

EFFECT: using the inventions allows to reduce measurement errors associated with sensor relative displacement.

14 cl, 9 dwg

Description

The invention relates to medical ultrasound diagnostic systems and, in particular, to ultrasound systems that perform measurements of tissue stiffness or elasticity using transverse (shear) waves.

Various tools have been developed to remotely study the mechanical properties of tissue for diagnostic purposes, which use the radiation power of an ultrasonic beam to remotely apply force to a region of tissue in the patient’s body (acoustic radiation force; also called “pushing” pulses). The strength of the acoustic radiation can be applied in such a way that it is possible to measure the elastic properties, or visually distinguish regions of different stiffness locally at the deformation point by tracking the deformation directly using further ultrasound imaging to follow the deformation model quasistatically. See, for example, Nightingale, K.R. et al, “On the feasibility of remote palpation using acoustic radiation force”, J. Acoust. Soc. Am., Vol. 110 no. 1 (2001), pp. 625-34; and M.L Palmieri et al. The deformation caused by the strength of acoustic radiation can also be used as a source of transverse waves propagating away from the deformation zone, which can then be displayed to study neighboring regions on their material properties by visualizing the velocity of the transverse waves in the time domain. See related publications by Sarvazyan, A. et al., “Shear wave elasticity imaging: A new ultrasonic technology of medical diagnostics”, Ultrasound Med. Biol. 24, pp 1419-1435 (1998) and ”Quantifying Hepatic Shear Modulus In Vivo Using Acoustic Radiation Force”, Ultrasound in Med. Biol., Vol. 34, 2008. This method can also be used to estimate the frequency domain, wave shear modulus, and viscosity. See Fatemi, M. et al., “Ultrasound-stimulated vibro-acoustic spectrography”, Science 280, pp 82-85 (1998). These methods use a one-dimensional array of emitters to generate shear waves and are thus associated with a limited effective penetration depth and a combination of weak coupling and safety restrictions imposed on the maximum power of the exciting rays, combined with adverse diffraction effects that limit the penetration depth for effective measurement. See Bouchard, R. et al., “Image Quality, Tissue Heating, and Frame Rate Trade-offs in Acoustic Radiation Force Impulse Imaging,” IEEE Trans. UFFC 56, pp 63-76 (2009).

In addition to this, existing methods, due to the limited degree of the pushing excitation pulse and the two-dimensional visualization methodology, are not able to distinguish between the areas of change of properties that are in the plane of visualization and those that can be nearby, but outside the plane. Mixing these property values outside the visualization plane with the values in the visualization plane during the visualization process can lead to an unnecessary decrease in the accuracy and diagnostic value at the output of these methods.

In conventional visualization and quantitative measurements using the acoustic radiation power currently practiced, the shock is generated by a one-dimensional array that produces a beam that can be easily controlled in one visualization plane, but limited to one, small magnitude, focus in the intersecting plane, or elevation plane , fixed focus mechanical lens. This leads to a mechanical push force that creates a lateral response in all directions, to and from the plane of the array. The movement of the fabric caused by this push propagates mainly radially in all lateral directions, and damps with a factor of 1 / R in the radial directions (in the case of a linear source in the direction of the pushing pulse) in addition to the normal damping caused by the viscosity of the fabric. In the case of qualitative and quantitative imaging using the strength of acoustic radiation, this is harmful, because areas of stiffness that are outside the plane will contribute to axial displacement in the image plane, reducing the accuracy of measuring stiffness in the image plane. In the case of quantitative measurement, the radial distribution distributes the useful energy of the shear wave outward from the image plane, reducing the amplitude of the signal necessary for an accurate assessment of the properties.

The movement produced by the transmission of acoustic radiation power within the permissible limits of the limitations on diagnostic radiation is very small, in the range from 0.1 to 15 microns in amplitude. Measurement of such small movements is carried out by tracking reflections from local inhomogeneities in the tissue under study, which means that the influence of a shear wave on the received signal can be difficult to distinguish. In addition to this, the movement of a shear wave is strongly attenuated in a tissue that is viscoelastic in nature. Thus, an adequate signal to noise ratio is difficult to obtain, and the penetration range is very limited. Any interfering signals adversely affect the results. An important source of interference is the relative movement of the sensor used to study and the region of tissue being studied. This can be caused by external sources, such as shaking the operator’s hand, or internal sources, such as breathing, heartbeat, or other voluntary or involuntary movements of the object. Attempts to improve the signal-to-noise ratio for methods of using the acoustic radiation power of the prior art have used bandpass signal filtering to eliminate low frequencies from the data. Most motion artifacts have a frequency below 50 Hz, so some improvement can be made. See, for example, Urban et al, “Error in Estimates of Tissue Material Properties from Shear Wave Dispersion Ultrasound Vibrometry,” IEEE Trans. UFFC, vol. 56, No. 4, (Apr. 2009). However, some of these interferences are quite large in amplitude, and band pass filtering is not always sufficient to eliminate adverse effects. Artifacts in the form of an incorrect estimate of displacements and, therefore, incorrectly calculated velocities and shear wave moduli are common.

Accordingly, it is an object of the present invention to improve the effective penetration depth of acoustic radiation effects, such as transverse (shear) waves. Another objective of the present invention is to reduce off-plane influences during material evaluation. Another objective of the present invention is to reduce the measurement error associated with the relative motion of the sensor in studies based on the strength of acoustic radiation.

In accordance with the principles of the present invention, a diagnostic ultrasound imaging system and a method are described that allow a user to obtain high resolution image data sufficient to measure tissue movement or the characteristics of a shear wave propagating through a tissue. An ultrasound probe with a two-dimensional array of transducer elements transmits a pushing pulse in the form of a wide band of energy into the tissue. A wide band of energy can be flat or non-flat, and can be obtained using a sequence of individually transmitted pulses of ultrasound or by transmitting a plane wave front. Unlike single vector pushing pulses from the prior art, a two-dimensional pushing pulse of a wide energy band produces a plane or half-plane shear wavefront that does not suffer from attenuation of energy propagation with a 1 / R coefficient inherent in the prior art. In accordance with yet another aspect of the present invention, a plurality of background tracking pulses are transmitted to an area around the location of the pushing pulse and the region of interest in which a transverse wave is to be detected. The echoes received from the background tracking pulses correlate over time to estimate the background motion in the region of interest during shear wave propagation, which is used to correct the measured displacement caused by the shear wave.

ON THE DRAWINGS:

FIG. 1 illustrates in block diagram form an ultrasound diagnostic imaging system made in accordance with the principles of the present invention.

FIG. 2a-2c illustrate the transmission of a sequence of pushing pulses to various depths to produce a shear wavefront.

FIG. 3 spatially illustrates the train of pulses along the pushing pulse vector, the resulting shear wavefront, as well as the series of tracking pulse vectors.

FIG. 4 illustrates the radial propagation of a shear wave front coming from a pushing pulse vector.

FIG. 5 illustrates a two-dimensional pushing pulse obtained in accordance with the principles of the present invention.

FIG. 6 illustrates a curved two-dimensional pushing pulse obtained in accordance with the principles of the present invention.

FIG. 7-9 illustrate the use of background tracking pulses to evaluate the background movement of tissue in a shear wave region in accordance with the principles of the present invention.

Turning first to FIG. 1, an ultrasound system constructed in accordance with the principles of the present invention for measuring shear waves is shown in block diagram form. The ultrasound probe 10 has a two-dimensional array 12 of transducer elements for transmitting and receiving ultrasonic signals. A two-dimensional array of transducers can scan a two-dimensional (2D) transmission plane by emitting beams and receiving returning echoes along one plane in the body, and can also be used to scan a volume region by emitting beams in different directions and / or planes of a volume (3D) region body. The array elements are connected to a beam former micro-shaper 38 located inside the probe, which controls the transmission of elements and processes the echo signals received from groups or subarrays of elements into signals with a partially formed radiation pattern. Signals with a partially formed radiation pattern are supplied from the probe to the multi-line transmitter 20 of the radiation pattern in the ultrasound system by the transmit / receive switch (T / R) 14. The coordination of the transmission and reception by the radiation conditioners is controlled by the controller 16 of the radiation conditioners connected to the multi-line radiation conditioner of the reception and to the transmission controller 18, which supplies the control signals to the micro-conditioner. The beam former controller responds to signals produced in response to a user manipulating a user control panel 40 to control the operation of the ultrasound system and its probe.

The multi-line reception beam former 20 creates a plurality of spatially different reception lines (A-lines) of echo signals during a single receive-transmit interval. Echo signals are processed for filtering, noise reduction, etc. the signal processor 22, and then stored in the memory 24 A-line. Time-varying A-line samples related to the same spatial position of the vector are connected to each other in an ensemble of echo signals related to a common point in the image field. The RF echoes of A-line sequential sampling of the same spatial vector are cross-correlated by the RF A-line cross-correlator 26 to obtain a sequence of tissue displacement samples for each reference point on the vector. Alternatively, the A-lines of the spatial vector can be Doppler processed to detect the movement of the shear wave along the vector, or other phase-sensitive methods, such as tracking speckles in the time domain, can be used. The wavefront peak detector 28 responds to detecting a shear wave displacement along the A-line vector to detect a shear wave displacement peak at each reference point on the A-line. In a preferred embodiment, this is done by plotting a curve, although cross-correlation and other interpolation methods can also be used if desired. The point in time at which the shear wave displacement peak is reached is noted with respect to the times of the same event at other A-line positions, all with respect to the total timing, and this information is fed to the wavefront velocity detector 30, which differentially calculates the shear velocity waves from times of peak displacement on adjacent A-lines. This speed information is fed to a speed display map 32, which indicates the shear wave velocity at spatially different points in the two-dimensional and three-dimensional regions of the image. The speed display card is connected to the image processing unit 34, which processes the speed map, preferably overlapping the anatomical ultrasound image of the tissue, for display on the image display 36.

FIG. 2a-2c illustrate the transmission of a sequence of focused push pulses with a high MI value (for example, MI 1.9 or less, so as to be within the diagnostic limits set by the FDA) along one direction of the vector to obtain a shear wavefront. Pulses with a high MI value and a long duration are used so that enough energy is transmitted to shift the tissue down along the transmission vector and cause the development of a shear wave. In FIG. 2a, the probe 10 on the skin surface 11 transmits the first pushing pulse 40 into the tissue with the beam profile 41a, 41b to a predetermined depth of focus, indicated by the shaded region 40. This pushing pulse will shift the tissue in focus downward, as a result of which the transverse wave front 42 waves will come out from the displaced tissue.

FIG. 2b illustrates a second pushing pulse 50 transmitted by the probe 10 along the same vector and focused at a greater depth indicated by the shaded region 50. This second pushing pulse 50 biases the tissue in focus, causing a shear wavefront 52 emanating outward from the biased tissue . Thus, both transverse wavefronts 42 and 52 propagate laterally through the tissue, with the initial wavefront 42 leading in front of the second wavefront depending on the time interval between the transmission of two pushing pulses and on the difference in propagation delay due to a change in the propagation distance to focus.

FIG. 2c illustrates the transmission by the probe 10 of the third pushing pulse 60 to a greater depth, which generates outward radiation of the shear wavefront 62. In FIG. 2c shows that the composite wavefront from three pushing pulses, indicated by the profile of the composite wavefront 42, 52 and 62, extends to a noticeable depth in the fabric, from the shallow depth of the first pushing pulse 40 to the largest depth of the third pushing pulse 60. This allows you to measure the transverse a wave at a considerable depth of fabric. In the implementation of the system shown in FIG. 1, such a sequence of pushing pulses can be used to detect the propagation of a shear wave to a depth of 6 cm, which is a suitable depth for visualizing and diagnosing masses in the mammary glands.

It should be borne in mind that more or fewer pushing pulses can be transmitted along the pushing pulse vector, including one pushing pulse. Multiple pushing pulses can be transmitted in any order, which actually determines the shape and direction of the composite wave front of the transverse wave. For example, if the push pulses shown in FIG. 2a-2c were transmitted sequentially from the deepest (60) to the smallest (40) with a corresponding delay between transmissions, the composite shear wave front shown in FIG. 2c will have a slope opposite to the slope shown in FIG. 2c. As a rule, each pushing pulse is a long pulse with a duration of 50 to 200 microseconds. A typical duration is, for example, 100 microseconds. An ultrasound generated by a pulse of 100 microseconds in duration is a pulse of a compression wave and can have a frequency of 7 or 8 MHz, for example. The push pulses are well focused, preferably with a number f from 1 to 2. In one typical implementation, the push pulse is transmitted every 2.5 ms (provided that the shear source travels from (40) to (50) and from (50) to ( 60) greater than the shear wave velocity), which gives push pulses with a transmission frequency of 400 Hz. In another implementation, all three pushing pulses are transmitted in the same sequence to trigger the full shear wavefront before the A-line tracking begins.

FIG. 3 is another illustration of the use of three push pulses to create a composite shear wavefront. Three pushing pulses are transmitted along vectors 44, 54 and 64, which, as can be seen, are aligned along one vector direction in FIG. 3. When the deepest pushing momentum of vector 64 is transmitted first, followed by pushing pulses focused at successively shallower depths, the shear wave fronts of the corresponding pushing pulses will propagate as indicated by waves 46, 56 and 66 in time shortly after being transmitted last pushing impulse (vector 64). The time intervals between the transmission times of the pushing pulses are determined by the shear and longitudinal velocities, because the propagation time to the focus must be taken into account. As the transverse waves 46, 56, and 66 move outward from the pushing pulse vector, they are tracked by tracking pulses 80, shown in spatial progression along the top of the drawing. Tracking pulses can occur both between and after pushing pulses. In contrast to the image of FIG. 2c, an illustration of the transverse waves 46, 56, and 56 of the composite wavefront in FIG. 3 shows that propagating transverse waves are substantially aligned in time and horizontal propagation distance. From the point of view of a sharp difference in the propagation speed between longitudinal pushing pulses and transverse waves in tissues, of the order of 100 to 1, a typical picture is when individual pushing pulses are transmitted in quick succession. Since the only function of the pushing pulses is the application of force to the tissue and no further period of time is required to receive the reflected signal, as is the case with echo pulse ultrasound imaging, after each pulse, essentially no delay time is required, and the pushing pulses can be transmitted in very quick succession. The travel time of the pushing pulse in the tissue is about 100 microseconds (ultrasound travels in the tissues at a speed of about 1560 m / s), while the travel time of the transverse wave in the tissue is about 2 to 10 milliseconds (the transverse waves travel in the tissues at a speed of about 1-5 m / s in tissue). Thus, from the point of view of the periodicity and speed of the shear wave, a fast sequence of pushing pulses is almost instantaneous, providing a single wavefront.

In conventional visualization and quantitative measurements using the strength of acoustic radiation, the pushing pulse (s) is transmitted along one direction of the vector. When the push is generated by a one-dimensional array, i.e. with a probe having one line of transducer elements, the array produces a beam that can be easily controlled in one plane of the image of the array, but is limited to one relatively small size focus in the transverse or longitudinal plane with the fixed focus of the mechanical lens of the probe. This leads to a mechanical shock, which creates a response emitted laterally in all directions, in and out of one plane of the image of the array. The tissue movement caused by the energy of this jolt propagates approximately radially in all lateral directions, as shown by circular wave fronts 72 surrounding the vector of the pushing pulse and outward directed arrows 70 in FIG. 4, and experiences attenuation of energy with a factor of 1 / R in radial directions in addition to the usual attenuation in tissue. In the case of qualitative and quantitative imaging using the strength of acoustic radiation, this is harmful, because non-plane areas of stiffness change will contribute to the axial displacement of tissues in the image plane, reducing the accuracy of measuring stiffness in the image plane. In the case of quantitative measurements using the strength of acoustic radiation, the radial propagation removes the useful shear wave energy from the image plane, decreasing the signal amplitude needed to evaluate the property.

In accordance with the principles of the present invention, a pushing pulse is formed as a two-dimensional wide band of energy instead of a single one-dimensional vector. Such two-dimensional wide push beams extend in the depth measurement D, as well as in the measurement E in height or azimuth, as shown by the wide push beam 80 in FIG. 5. A wide push beam 80 leads to the formation of transverse waves with plane fronts, as shown by plane wave fronts 90, 92 in FIG. 5, which move laterally from the force field of the wide push beam 80, as shown by arrows 91, 93. This exciting transverse wave is a plane wave source and not the linear source shown in FIG. 4, eliminating 1 / R attenuation in radial energy dissipation. The programmability and flexibility of the two-dimensional array 12 for the formation of beams in arbitrary directions and from obvious centers at various places on the surface of the array are used to form pushed tissue areas of the general shape, size and direction of the push by axial and / or lateral swing of the focal spot, quickly jumping the focal spot from one place to another, or both, taking advantage of the substantial relationship between the propagation velocities of longitudinal pushing waves and transverse flax in the tissue (of the order of 100 to 1) in order to ensure the formation of an effective source of shear waves, which can have a fairly arbitrary size, shape and orientation, so that a focused and controlled source of a two-dimensional or three-dimensional shear wave beam of the desired orientation, shape and size can be formed .

In a simple implementation of the present invention shown in FIG. 5, a flat, widened strip of pushed fabric 80 is excited, which forms transverse waves in a flat strip 90, 93, propagating laterally, rather than radially outward, and damping with increasing distance traveled by the transverse wave. This improves the penetration distance of various modalities of the radiation power. This layer can be formed by focusing deep in the tissue and starting the transmission of a long ultrasound packet. While the packet is being transmitted, the focal point is shifted to a smaller position, towards the probe, to form a linear source. Several such lines of the force push beam are transmitted in a plane perpendicular to the surface of the probe, as shown in FIG. 5. Alternatively, the plane of the force of the push beam is transmitted in other planes not perpendicular to the surface of the array and within the direction of the array to create a flat source of shear waves. Such a transmission will effectively produce a single pushing force in two dimensions, for such a time that the duration of the entire exciting sequence is somewhat faster than the period of the received shear waves. Since the longitudinal propagation time of ultrasound is about 100 μs, while the desired shear wave period is about 2 ms, there is time for multiple transmissions to create an energy band 90, 92.

A variation of the transmission method depicted in FIG. 5 is a transmission in a strip beam that, by simultaneously exciting the elements of a two-dimensional array of transducers in height and azimuth, transmits a strip beam from a two-dimensional array. Since the delay profile of a two-dimensional array is fully programmable, transmitting a strip focused deep in the field, and then moving the focal point closer with a speed comparable to the velocity of the transverse wave, will provide the formation of a simple plane source of the transverse wave. This flat source can be transmitted at any angle of rotation, so that transverse waves can propagate in any lateral direction. In addition, the angle of inclination of the planar source can vary, so that the shear wave source can be directed in a plane that is not perpendicular to the array.

A third implementation of the present invention is illustrated in FIG. 6. In this embodiment, the strip beam is transmitted by a two-dimensional array of transducers, which is curved in the transverse direction, either in space or in the delay profile, or both, so that the resulting shear wave source is focused into a thin beam, which further increases the resolution and sensitivity of the methods used to detect it. It is even possible to create curvature in the axial direction, as shown by the wide push beam (PBS) in FIG. 6, creating a two-dimensional focusing of transverse waves. As this drawing shows, a two-dimensional 12 array of transducers forms a curved wide push beam PBS. The curvature of the PBS causes the SWF shear wave front to gradually converge as it passes, as shown by the gradual convergence of SWF1, SWF2, and SWF3 to the cell 98-shaded plane. This convergence is also shown by the profile 96 of the curved shear wave fronts. The shear wave front SWF2 is shown on the right in the drawing, illustrating the inverse curvature of the shear wave front as it passes beyond the maximum convergence line of SWF3. This method of focusing shear waves is most suitable for a linear measurement method, and not for a planar measurement method. The speed of data collection decreases sharply in exchange for a significant increase in sensitivity near the focus of the shear wave in SWF3. This method can also be used to focus a two-dimensional curvilinear wave front of a transverse wave to a focal point limited only by diffraction or to a region with a limited axial depth.

Diagnostics of tissue stiffness, which is performed by measuring shear waves, is highly dependent on accurately tracking the shear wavefront in time so that changes in its propagation velocity as it passes through different tissues can be accurately measured. In prior art systems, these measurements were carried out under the assumption that there was no relative movement between the ultrasound probe and the tissue, so that the only relative movement of the tissue is the movement produced by the force of the pushing pulse. This assumption is often incorrect, since relative movement can also come from unsteady retention of the probe, from patient movement, or from anatomical movement due to respiration and heartbeat. The bias caused by the radiation force is very small, on the order of 10 microns. Although the accuracy of ultrasonic RF tracking can reach 1–2 μm, shear wave motion can be blocked by much larger patient movements, such as cardiac and respiratory movements, as well as ambient noise. Although filtering may be used to try to eliminate noise whose frequency is outside the harmonic frequency range of the shear wave, in accordance with another aspect of the present invention, an additional step has been taken to reduce noise. This step consists in using an offset estimated far from the field of excitation (for example, at a depth of at least half the depth of field, which is far from the focus in the depth direction) as background noise, since it can be assumed that no no significant radiation power. This “source” of noise in the form of an offset estimate is subtracted from the shear wave offset estimated in the region of interest.

A simple example of background motion detection is shown in FIG. 7. A single-vector pushing beam, being focused, has the most significant effect along the axis of the beam near a depth of 110 focal points. FIG. 7 illustrates a profile 100 of a vector push beam in which the force of the push beam is concentrated. Some elastographic methods based on the strength of acoustic radiation include only tracking along the same axis as the push beam, in which case data from tracking beams that are already in use can be used to detect background motion, but from ranges which are significantly shorter and significantly longer than the focal length, that is, outside the depth of field of the focused pushing beam, to make an axial motion estimate to subtract from the measurement reference. Asterisks 102 and 104 illustrate the focus areas of two background tracking beams, one located above the focus area of the pushing pulse and the other below the focus area of the pushing pulse. The focus areas of the background tracking beams are shown by dashed beam profiles on both sides of the background tracking beams. Echoes from these background tracking locations are digitized multiple times before, during and / or after the transmission of the pushing pulse (s). These time-varying echoes are compared, usually by correlation, and comparisons are used to evaluate the presence of axial background motion. Any tissue displacement due to background effects is subtracted from the estimates of the movement caused by the shear wave to adjust the estimate of the movement of the shear wave by the amount of background effects.

FIG. 8 illustrates another example of background motion detection. In this example, additional points 106, 107, 108 and 116, 117, 118, located laterally far beyond the shear wave tracking region 120 of interest, can be monitored during the measurement interval to obtain data that allows you to calculate an estimate of background motion at any point in or around the area of interest. In this way, motion effects from tilting or rotating the probe during the measurement interval can be detected. For example, if comparisons of changes in echoes at points 106, 107, 108 over time indicate movement up and to the left of the region of interest 120, and differences at points 116, 117, 118 at the same time indicate movement down and to the right of of the region of interest, it can be concluded that there is a general rotational or oblique movement of the probe with respect to the region of interest for which compensation of the measurements should be made.

As shown in FIG. 9, it is also possible, and in some cases it may be desirable, to track in the plane of a two-dimensional image along several lines in the region of interest 120 adjacent to the shock region 100 at times before and after the shock event. In this example, two background motion tracking lines 126, 127, 128 and 136, 137, 138, are periodically digitized during the measurement interval to the left of the pushing pulse vector 100 and two background motion tracking lines 146, 147, 148 and 156, 157, 158 are digitized on the right from the vector of the pushing momentum. Typically, you need one sample of background motion values before and several samples of background motion values after the push event to get an estimate of the motion due to the push event. However, if two or more ensembles of samples of the background motion values are obtained before the push, an estimate of the background motion can also be obtained. If at least one ensemble of background motion echoes is also obtained a sufficiently long time after the push, additional estimates of the background motion can also be obtained, since the motion can be interpolated, not extrapolated, in time from the moment before the pushing pulse to the moment after pushing momentum. This method can be performed with multiple lateral displacements from the axis of the pushing beam. If the background motion is not uniform in the region of interest, an estimate of the scalar field of the axial component of the motion within the volume of the sample can be obtained.

It should be borne in mind that the correction taking into account the background motion can be performed for measurements made in three-dimensional space, in addition to one plane. Using a two-dimensional array of transducers, as shown in FIG. 1, allows the use of the effectiveness of three-dimensional elastography for additional clinical benefits, since out-of-plane changes in elastic properties can adversely affect the effectiveness of uniplanar elastographic measurements. Additional three-dimensional control of the geometry of the push beam by means of a two-dimensional array can improve the signal-to-noise ratio and provide additional functionality. In this case, additional background motion tracking beams outside the region of interest during the measurement interval and / or early and late background motion tracking beams can be added in the three-dimensional region of interest, as indicated above for the two-dimensional case, to obtain full three-dimensional volumetric axial motion estimation to correct the measured response to excitation from the push beam. For example, four background motion tracking lines can be transmitted at 90 ° intervals around the pushing pulse vector. Background movement tracking lines can be transmitted in front and behind a two-dimensional band of a pushing pulse, as described above, in order to detect tissue movement in a three-dimensional space above which transverse waves are measured.

Claims (14)

1. Ultrasonic diagnostic imaging system for transverse-wave analysis, containing:
an ultrasonic array probe containing a two-dimensional array of transducer elements that transmits a pushing pulse along a given vector to generate a shear wave, transmits tracking pulses along tracking lines adjacent to the pushing pulse vector, and receives echo signals from points along the tracking lines, the focused pushing pulse It has a measurement both in height and in azimuth;
memory for storing echo data of the tracking line;
a motion detector responsive to the tracking line data for detecting a shear wave passing through the locations of the tracking lines; and
a display for displaying the characteristics of the detected shear wave,
moreover, the ultrasonic matrix probe is additionally configured to transmit along one or more background tracking lines, along which background motion echoes are received, in the vicinity of the pushing pulse vector at various times, which are compared to detect background motion near the transverse wave. 2. The ultrasound diagnostic imaging system according to claim 1, wherein the background motion signals received at various points in time are compared by correlation processing. 3. The ultrasound diagnostic imaging system according to claim 1, in which the line tracking background motion is located along the vector of the pushing pulse,
moreover, background motion is detected at points located above and below the focus depth of the pushing pulse. 4. The ultrasound diagnostic imaging system according to claim 1, wherein the background movement tracking line is located on either side of the pushing pulse vector. 5. The ultrasound diagnostic imaging system according to claim 1, further comprising a region of interest for shear wave analysis located on at least one side of the pushing pulse vector,
moreover, the tracking line of the background movement is located next to the area of interest. 6. The ultrasound diagnostic imaging system according to claim 1, further comprising a region of interest for shear wave analysis located at least on one side of the pushing pulse vector,
wherein the background motion tracking line is located in the region of interest. 7. The ultrasound diagnostic imaging system according to claim 6, wherein the second background tracking line is located in a region of interest adjacent to the first background tracking line. 8. The ultrasound diagnostic imaging system according to claim 1, in which the ultrasonic matrix probe further comprises a two-dimensional array of transducer elements,
moreover, the background motion signals are located in three-dimensional space around the vector of the pushing pulse. 9. The ultrasound diagnostic imaging system of claim 8, wherein the background motion signals are received from volumetric quadrants located around the vector of the pushing pulse. 10. The ultrasound diagnostic imaging system according to claim 8, in which the two-dimensional matrix transducer element is further configured to produce a push-pulse energy band to form a shear wave front,
wherein, the background motion signals are received from areas in front and behind the shear wave front. 11. The ultrasound diagnostic imaging system according to claim 1, further comprising a processor responsive to the measurement of the shear wave and the detected background motion, which is configured to correct the measurement of the shear wave based on the detected background motion. 12. A method for correcting the measurement of a shear wave characteristic measured in a tissue region, taking into account the effect of relative motion between the tissue and the ultrasound probe, comprising the steps of:
using an ultrasound probe to transmit a focused pushing pulse to generate a shear wave in the tissue region, the ultrasound probe containing a two-dimensional array of transducer elements;
transmit along one or more background motion tracking lines and receive background motion echoes along the background motion tracking lines outside the focus depth of the focused pushing pulse;
comparing the background motion echoes received at different times to detect the relative motion between the ultrasound probe and the tissue region during the time interval when a shear wave is detected, and
form the characteristic of the shear wave, which is adjusted for relative motion. 13. The method according to p. 12, in which the use of an ultrasonic probe further comprises the step of transmitting a pushing pulse by means of an ultrasonic probe for generating a transverse wave,
moreover, the detection of relative motion further comprises stages in which:
receive echoes of the background motion near the pushing pulse at different points in time and
echoes of the background motion obtained at different times are compared to detect background motion. 14. The method of claim 13, wherein comparing the background motion echoes received at various times, further comprises the step of correlating the background motion echoes.

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